(1) Field of the Invention
The present invention relates to magnetic resonance imaging (MRI) methods, image display, and image processing for delineation of tissue structures and interfaces of the structures. More particularly, the invention relates to a method for measuring (a) diffusion gradient in tissues and (b) diffusion anisotropy and diffusion gradient simultaneously in tissues.
(2) Description of the Related Art
Any nucleus with a non-zero magnetic moment such as proton in water molecule tends to align itself with the direction of a magnetic field when placed in the field. The nucleus precesses around the field direction at a characteristic angular frequency (Larmor frequency ω) which is dependent on the strength of the magnetic field (B0) and on the properties of the specific nuclear species (the gyromagnetic ratio γ), e.g., ω=γ B0. Nuclei exhibiting such phenomena are referred to as spins.
In the presence of a static uniform magnetic field (polarizing field B0), each individual magnetic moment of the spins in the tissue tends to align with the polarizing field, yielding a net magnetization parallel to the field (longitudinal magnetization). In the transverse plane to the field, however, the randomly oriented magnetic components of the spins cancel one another, producing no net magnetization perpendicular to the field.
By applying a transient magnetic field (excitation field B1), the longitudinal magnetization may be rotated, or tipped, away from the polarizing field and the transverse component of the tipped magnetization precesses around the polarizing field at the Larmor frequency. The degree to which the longitudinal magnetization is tipped, and hence, the magnitude of the transverse magnetization depends primarily on the length of time and magnitude of the applied excitation field B1. When the excitation field is terminated, the recovery of the longitudinal magnetization to its equilibrium value along the polarizing field is described by the spin-lattice relaxation process that is characterized by the time constant T1. This recovery process is also referred to as the longitudinal relaxation process.
Once the excitation field is removed, precession of the transverse magnetization around the polarizing field may induce a sinusoidal signal in a receiving coil, and the frequency of the signal is the Larmor frequency. The net magnetization gradually reorients itself with the polarizing field when the excitation field is removed, resulting in the amplitude of the signal decaying exponentially with time. The rate at which the signal decays depends on the homogeneity of the magnetic field and on T2, which is referred to as the “spin-spin relaxation” constant, or the “transverse relaxation” constant. The T2 constant is inversely proportional to the exponential rate at which the aligned precession of the spins would dephase after removal of the excitation field in a perfectly homogeneous field.
In order to produce a spatial image of a substance such as human tissue, magnetic field gradients are employed to obtain MRI signals from specific locations in the substance. Although the direction of the polarizing field remains the same, its strength varies along the x, y, and z axes oriented with respect to the substance. Varying the strength of the polarizing field linearly along one axis causes the Larmor frequency to vary linearly as a function of its position along the axis. Thus, by controlling the strength of these gradients during each MRI cycle, the spatial distribution of spin excitation can be controlled and the location of the MR signals can be identified. Then, the received spatial-encoded MR signals are digitized and processed to reconstruct the MR image using one of many well known reconstruction techniques.
MRI is a non-invasive modality for imaging the internal structures of the human body with no known health hazards attributed to its use. It is a routine clinical diagnostic imaging tool.
Diffusion-weighted MRI combines magnetic resonance imaging principles with the effects of molecular diffusion on MR signal in order to reveal diffusion characteristics of water molecules in tissues (Le Bihan, D. et al (1986). MR Imaging of Intravoxel Incoherent Motions: Application to Diffusion and Perfusion in Neurologic Disorders. Radiology 161:401-407; Merboldt, K. D., et al (1985). Self-diffusion NMR imaging using stimulated echoes. J. Magn. Reson. 64:479-486; and Taylor, D. G. and Bushell, M. C. (1985). The spatial mapping of translational diffusion coefficients by the NMR imaging technique. Phys. Med Biol. 30, 345-349)). Molecular diffusion refers to the random Brownian motion of molecules resulting from thermal agitation. During their random, diffusion-driven displacements molecules encounter various tissue components, i.e., macromolecules, intracellular organs, cell membranes, fibers, etc. These micrometer-sized cellular structures impede the molecular displacements. This diffusion effect of molecular displacements, due to the presence of the tissue microstructure, is encoded in the MR signal by using magnetic field gradient pulses (diffusion-encoding gradients). The observation of this displacement distribution provides unique information regarding the structure and geometric organization of tissues (Le Bihan, D. et al (1986). MR Imaging of Intravoxel Incoherent Motions: Applications to Diffusion and Perfusion in Neurologic Disorders. Radiology 161:401-407).
In diffusion-weighted MRI, the signal contrast is based on spin dephasing resulting in water molecule diffusion, and the rate of diffusion is measured by applying diffusion-encoding gradients. In comparison to free water diffusion, the diffusive motion of water molecules in tissue is influenced by the tissue microstructure. The degree of restriction that is encountered by water diffusion in a tissue is reflected in the diffusion coefficient or diffusivity, i.e., the stronger the restriction, the smaller the diffusivity. Water diffusion causes MR signal attenuation and provides a means for direct measurement of the diffusion coefficient, resulting in diffusion-weighted MRI (DWI).
Water diffusion-induced signal loss in a DWI image voxel results from the total contribution of all water molecules within the voxel. The overall effect observed in the voxel reflects, on a statistical basis, the displacement distribution of the water molecules present within this voxel. Thus, the measured diffusion coefficient is an apparent diffusion coefficient (ADC). This diffusion constant is estimated from the linear relationship between the logarithm of the echo intensity and the magnitude of the diffusion-encoding gradient in which ADC appears as a constant of proportionality (Stejskal, E. O. Use of Spin Echoes in a Pulsed Magnetic-Field Gradient to Study Anisotropic, Restricted Diffusion and Flow. J. Chem. Phys. 43:3597-3603).
The level of tissue microstructure restriction on water diffusion is reflected in relative changes in the ADC. In an acute stroke, the ADC in the ischemic territory was reduced by as much as 50% within minutes after the onset of ischemia, demonstrating the value of DWI for the early detection of stroke (Moseley, M. E., et al (1990). Diffusion-weighted MR imaging of acute stroke: correlation with T2-weighted and magnetic susceptibility-enhanced MR imaging in cats. AJNR 11, 423-429). The visualization of changes in the diffusion properties of tissue water with DWI has become a routine clinical tool to characterize tissue structure and to identify and differentiate disease processes.
Molecular diffusion is a three-dimensional process. In pure liquids such as water, individual molecules are in constant motion in every possible direction and have an equal probability of moving in every direction. Therefore, diffusion in pure liquids is isotropic. Water diffusion in biologic tissue along one particular direction, however, as selected by the direction of the probing magnetic field gradient, could be different from the diffusion along another direction due to the presence of various tissue components that restrict the Brownian motion in certain directions, rendering diffusion in structured tissues anisotropy. As diffusion is encoded in the MRI signal by using diffusion-encoding gradients, only molecular displacements that occur along the direction of the diffusion-encoding gradients are visible. Thus, diffusion anisotropy in structured tissues can be examined by observing variations in the diffusion measurements when the direction of the diffusion-encoding gradients is changed.
Diffusion anisotropy in structured tissues is reflected in the measured direction-dependent ADCs. Since these ADCs depend on the direction of the diffusion-encoding gradients and are exquisitely sensitive to the choice of the laboratory frame of reference, they provide only a qualitative description to diffusion anisotropy. To overcome this deficiency, a diffusion tensor is introduced to replace the diffusion coefficient (Basser, P J, et al. (1994). Estimation of the Effective Self-Diffusion Tensor from the NMR Spin Echo. Journal of Magnetic Resonance, Series B 103, 247-254). This technique is referred to as diffusion tensor MRI (DTI). With DTI, diffusion anisotropy effects can be fully extracted, characterized, and exploited. It provides quantitative parameters to characterize intrinsic features of tissue microstructure objectively. These parameters are independent of the direction of the diffusion-encoding gradients and of the choice of the laboratory frame of reference. A direct application of the DTI technique is the development of white matter fiber tractography (Basser, P J, et al. (2000). In Vivo Fiber Tractography Using DT-MRI Data. Magn Reson Med, 44: 625-632).
In summary, both DWI and DTI investigate water molecule diffusion with MRI. DWI examines the effect of diffusion on MRI signal in the direction of diffusion-encoding magnetic field gradients. DTI studies anisotropic diffusion, i.e., spatial-direction varied diffusion. DWI is a routine clinical tool. DTI has been invented in 1993 and has been widely used in both research and clinical today.
The diffusion tensor obtained with DTI determines the spatial-direction variation of diffusion within an MR image voxel, but it provides no information about intravoxel spatial-location variation of the diffusion. Inhomogeneous tissue structures yield diffusion varying from location to location. Intravoxel spatial-location related diffusion (diffusion gradient) has not been explored yet, neither theoretical nor experimentally. Within an MRI voxel, when diffusion is faster in one location than another location, it exhibits a spatial diffusion gradient between the two locations. For example, when diffusion is slower in the left half of a voxel than that in the right half, it produces a spatial gradient from left to right within the voxel. This diffusion gradient is characterized by the spatial derivative of diffusion coefficient. Since these diffusion gradients are caused by the inhomogeneous tissue structures, a measurement of the diffusion gradients with MRI will offer a noninvasive imaging tool for investigating intravoxel inhomogeneous tissue structures. In biologic tissues including very structural ones, intravoxel diffusion gradient in most places is not expected to be large except in the boundary of two structures having markedly different diffusion properties. The large gradient on the tissue interface and small gradients within the tissues could make a sharp contrast of the interface from the tissues, providing a quantity to delineate the interface precisely. Thus, if the diffusion gradient is measurable, it provides valuable information to elucidate intravoxel spatial-location variation of tissue structures.
The patent (U.S. Pat. No. 6,163,152 to Bernstein et al) presented a system and method for correcting systematic errors that occur in MR images due to magnetic gradient non-uniformity. In other words, the method improves the quality of MR images. The method has no relation with diffusion MRI though it can be applied to improve the quality of diffusion MR images. The application (U.S. application Ser. No. 2006/0001424 to Harvey et al) presented a method for MR imaging of at least a portion of a body placed in a stationary and substantially homogeneous main magnetic field. This method is related with DWI. The application (U.S. application Ser. No. 2004/0227510 to Rose et al) presented a method for eddy current compensated diffusion imaging. It corrects the eddy current artifact in DWI due to diffusion-encoding magnetic field gradient pulses. The application (U.S. application Ser. No. 2004/0113615 to Bammer et al) presented a mathematical framework for characterizing and then correcting the contribution of gradient non-uniformities to diffusion tensor imaging (DTI). In other words, the method improves the quality of DTI images. The application (U.S. application Ser. No. 2004/0140803 to Deimling) presented a specific pulse sequence (DESS) for DWI. The application (U.S. application Ser. No. 2003/0160612 to Yablonskiy et al) presented a specific method for DTI. The patents and applications are incorporated herein by reference in their entireties.